So-called biosensors are used for the determination of antibodies, antigens or DNA in blood, water, air or food. They are based on specific binding of the analyte to a capture molecule, for example of an antibody to an antigen or of a DNA sequence to the complementary DNA capture sequence. This binding event is often detected by optical methods. Fluorescent dyes are in this case introduced into the analyte, for example by PCR in the case of DNA analysis, and subsequently read out at the positions with different capture sequences. Such optical systems, however, are expensive and elaborate to handle. The required apparatus are sensitive and not suitable for field tests.
Electrochemical biosensors are a suitable alternative in this case.
Here, detection is carried out by the conversion of an electrochemical substance. Electrochemical detection is widespread, for example, in glucose sensors. The glucose is oxidized by a redox enzyme, namely glucose oxidase, and oxygen which is present is simultaneously reduced. The resulting hydrogen peroxide is then electrochemically oxidized again, and the glucose concentration is thus determined amperometrically. In refinements of these glucose sensors, mediators replace the oxygen. During the oxidation of glucose by the redox enzyme, the mediator, for example 1,1-dimethyl ferrocene, is simultaneously reduced. The electrochemical oxidation of the enzymatically reduced mediator, for amperometric determination of the glucose concentration, can take place at much lower potentials in this case, so that the measurement is more accurate and less susceptible to interference. In this context, many pulsed amperometric methods have been developed which determine not the total amount of redox mediator, but only that fraction which was previously reduced enzymatically.
For the electrochemical detection of binding events, however, the absolute quantity of a redox-active substance must be measured as sensitively as possible. The actual marker for the presence of the analyte at a sensor position is the enzyme per se. For DNA analysis, for example, a biotin marker is attached to the analyte during PCR. Various DNA capture sequences are bound to the different sensor positions on a biosensor. The analyte hybridizes only with the matching sequences, and the unbound analyte is washed away. A streptavidin-enzyme, for example alkaline phosphatase, is then bound to the biotin marker molecule in order to detect this binding event. If enzyme substrate is added, for example p-aminophenyl phosphate, then p-aminophenol is released by hydrolytic cleavage of the phosphate only at the sensor positions to which the analyte has bound.
The use of p-aminophenyl phosphate as a substrate for alkaline phosphatase is introduced in CLIN. CHEM 36/11, pp. 1941-1944 (1990), where a bead-based immunoassay is described. Here, alkaline phosphatase reveals whether analyte has bound to the bead-immobilized antibodies. The beads are incubated with p-aminophenyl phosphate and then the supernatant solution is examined for p-aminophenyl phosphate in a flow-injection analysis system. This solution flows through an electrochemical sensor, the working electrode of which is constantly polarized at about +0.1 V vs. Ag/AgCl reference electrode. Measurement in the flowing electrolyte or sample volume has the advantage that no significant depletion of the p-aminophenol takes place in front of the electrode. It is continuously replenished by the flow. Such a system, however, is not microsystem-compatible. The electrochemical sensor can read only one sample at a time, and the volumes required are large.
If an array of sensors with different capturers—whether DNA capture sequences or antibodies—is intended to be used, then the electrochemical detection must be carried out in a stationary electrolyte so that the sensors only detect the signal of the capturers immobilized directly thereon. The presence of the enzyme is revealed by a rise in the p-aminophenol concentration. Simple electrochemical sensors, in which the working electrode is at a constant potential, are not suitable for this. Owing to the conversion of the enzyme product taking place continuously, it will be consumed. A decrease in the concentration due to the measurement per se will thus be superimposed on the rise in the concentration due to the enzyme activity.
In order to circumvent this problem, U.S. Pat. No. 6,682,648 B1 proposes the use of interdigital electrode arrays. Each sensor consists of two interdigital electrodes. By a bipotentiostat, one of the electrodes is polarized positively and the other negatively. The p-aminophenol is oxidized at the first electrode and therefore consumed. If it can now diffuse to the second electrode, then it will be reduced again there and is once more available for the measurement at the first electrode.
A prerequisite for the latter redox cycling system is that the distance between the two electrodes should be very small, however, so that the transport of p-aminophenol and the oxidation product quinone imine by diffusion between the electrodes takes place rapidly enough, in respect of which reference may be made to the publication K. Aoki, J. Electroanal. Chem. 270 (1989), p. 35. In the aforementioned U.S. Pat. No. 6,682,648 B1, the use of interdigital electrodes with structure dimensions smaller than 1 μm is proposed for this purpose. The result of this is that the production of a biosensor array with such interdigital electrodes is elaborate and expensive.
Further information about the measurement, especially in liquids, or for biochemical measurements is given in DE 43 35 241 A1, DE 41 31 731 A1, DE 197 17 809 U1 and DE 199 17 052 A1. A method for the electrochemical measurement of redox cycling with a practicable electrode arrangement is described in detail in WO 01/67587 A1.